Implantable stimulation devices are devices that generate and deliver electrical stimuli to body nerves and tissues for the therapy of various biological disorders, such as pacemakers to treat cardiac arrhythmia, defibrillators to treat cardiac fibrillation, cochlear stimulators to treat deafness, retinal stimulators to treat blindness, muscle stimulators to produce coordinated limb movement, spinal cord stimulators to treat chronic pain, cortical and deep brain stimulators to treat motor and psychological disorders, and other neural stimulators to treat urinary incontinence, sleep apnea, shoulder sublaxation, etc. The description that follows will generally focus on the use of the invention within a Spinal Cord Stimulation (SCS) system, such as that disclosed in U.S. Pat. No. 6,516,227. However, the present invention may find applicability in any implantable medical device system. For example, the disclosed invention can also be used with a Bion™ implantable stimulator, such as is shown in U.S. Patent Publication 2007/0097719, filed Nov. 3, 2005, or with other implantable medical devices.
As shown in FIGS. 1A and 1B, a SCS system typically includes an Implantable Pulse Generator (IPG) 100, which includes a biocompatible device case 30 formed of titanium for example. The case 30 typically holds the circuitry and battery 26 necessary for the IPG to function, although IPGs can also be powered via external RF energy and without a battery. The IPG 100 is coupled to electrodes 106 via one or more electrode leads (two such leads 102 and 104 are shown), such that the electrodes 106 form an electrode array 110. The electrodes 106 are carried on a flexible body 108, which also houses the individual signal wires 112 and 114 coupled to each electrode. In the illustrated embodiment, there are eight electrodes on lead 102, labeled E1-E8, and eight electrodes on lead 104, labeled E9-E16, although the number of leads and electrodes is application specific and therefore can vary. The leads 102, 104 couple to the IPG 100 using lead connectors 38a and 38b, which are fixed in a header material 36, which can comprise an epoxy for example.
As shown in FIG. 2, the IPG 100 typically includes an electronic substrate assembly 14 including a printed circuit board (PCB) 16, along with various electronic components 20, such as microprocessors, integrated circuits, and capacitors mounted to the PCB 16. Three coils are generally present in the IPG 100: a telemetry coil 13 used to transmit/receive data to/from an external controller 12; a charging coil 18 for charging or recharging the IPG's battery 26 using an external charger (not shown); and a coil 66 (not shown in FIG. 2) used in the boost converter 150 used to generate a high compliance voltage, as discussed below in conjunction with FIG. 3. The telemetry coil 13 can be mounted within the header 36 of the IPG 100 as shown, or it can be mounted on the printed circuit board within the IPG.
As just noted, an external controller 12, such as a hand-held programmer or a clinician's programmer, is used to send data to and receive data from the IPG 100. For example, the external controller 12 can send programming data to the IPG 100 to dictate the therapy the IPG 100 will provide to the patient. Also, the external controller 12 can act as a receiver of data from the IPG 100, such as various data reporting on the IPG's status. The external controller 12, like the IPG 100, also contains a PCB 70 on which electronic components 72 are placed to control operation of the external controller 12. A user interface 74 similar to that used for a computer, cell phone, or other hand held electronic device, and including touchable buttons and a display for example, allows a patient or clinician to operate the external controller 12.
Wireless data transfer between the IPG 100 and the external controller 12 takes place via magnetic inductive coupling. To implement such functionality, both the IPG 100 and the external controller 12 have telemetry coils 13 and 17. Either coil can act as the transmitter or the receiver, thus allowing for two-way communication between the two devices, as explained further below. When data is to be sent between the external controller 12 and the IPG 100, the transmitting coil 17 or 13 is energized with alternating current (AC), which generates a magnetic field 29, which in turn induces a current in the other of coils 17 or 13. The generated magnetic field 29 is typically modulated using a communication protocol, such as a Frequency Shift Keying (FSK) protocol, which is well known in the art. The power used to energize the coil 17 or 13 can come from batteries 76 and 26 within the external controller 12 and IPG 100 respectively. The induced current in the receiving coil can then be demodulated back into the data signals that were transmitted.
Inductive transmission of data can occur transcutaneously, i.e., through the patient's tissue 25, making it particular useful in a medical implantable device system. During the transmission of data, the coils 13 and 17 preferably lie along a common axis in planes that are parallel. Such an orientation between the coils will generally improve the coupling between them, but deviation from ideal orientations can still result in reliable data transfer.
As shown in FIG. 3, a therapeutic current, lout, to be provided at a given electrode 106 (only one electrode in shown in FIG. 3 for convenience) is provided by a current source. In the illustrated example, the current source is digitally programmable and is referred to as a Digital-to-Analog Converter, or “DAC” 60. The current is provided to the patent's tissue, R, and is set with respect to a reference potential (e.g., ground) as designated generically by node 107, which may comprise another electrode 106, the case 30 of the IPG 100, etc. The electrode 107 may or may not be coupled to a DAC of its own. For example, if electrode 106 sources lout, electrode 107 may be programmed to sink lout to ensure that no charge builds up in the patient's tissue, R.
For the DAC 60 to be able to provide the desired output current, Iout, the DAC 60 must receive a power supply voltage, called the compliance voltage, V+, and which is generated by a boost converter 150. The boost converter 150 comprises one type of DC-DC conversion circuit and is used to convert the battery voltage, Vbat, to the compliance voltage V+. The compliance voltage V+ provides power to the electrodes or other loads in a more generic implantable medical device. The boost converter 150 is needed in an IPG 100 because the compliance voltage, V+, required to provide the desired therapeutic current, Iout, at the electrode may be higher than the battery voltage, Vbat. For example, the battery voltage, Vbat, may be in the neighborhood of 4V, while compliance voltages of 18-20V may be necessary to provide higher-magnitude therapeutic currents.
The compliance voltage V+ is adjustable depending on the power it must provide at any given time. Its optimal value at any given time depends on the magnitude of the programmed stimulation current, the resistance of the tissue R, and other factors. Adjustment of V+ is important in the IPG: if V+ is too low, the DAC 60 will become “loaded” and unable to provide the desired current, Tout; if V+ is too high, the DAC 60 will be able to provide the desired current, Iout, but battery power will be wasted, because some portion of the compliance voltage V+ will be dropped across the DAC 60 without any useful effect.
Adjustment of V+ is made by V+ monitor and adjust logic circuitry 62, which determines whether V+ needs to be raised or lowered via a feedback loop. V+ monitor and adjust logic circuitry 62 can comprise part of the IPG's microcontroller 155 (see FIG. 4), or may be a standalone circuit block. If V+ is too low, circuitry 62 outputs a “boost” signal to a pulse width modulator 63. The pulse width modulator adjusts the pulse width of a clock signal, CLK, in a manner specified by a pulse width, PW, provided by the IPG's microcontroller 155 (FIG. 4). The pulse-width-modulated pulse train is sent to the gate of a transistor 64. When the transistor 64 is on, current passes through an inductor 66, which can comprise a dedicated inductor used exclusively in the boost converter 150, or can comprise one of the coils 18 or 13′ (FIG. 2) in the IPG 100. Later, when the transistor 64 is turned off, the current in the inductor 66 must discharge and does so through diode 68 to charging capacitor 69, whose top plate comprises the compliance voltage V+. Because the capacitor 69 was already charged to the battery voltage, Vbat, the additional charge from the inductor 66 boosts the compliance voltage V+ to a value higher than Vbat. Diode 68 prevents this excess charge from dissipating backwards into the circuit, and the capacitor 69, in addition to storing the charge, also filters the compliance voltage to stabilize it. Thus, as the gate of transistor 64 oscillates between on and off, the compliance voltage V+ continues to boost. If V+ monitor and adjust logic circuitry 62 determines that V+ is too high, it disables the “boost” signal. This halts oscillations at the gate of the transistor 64, which causes V+ to fall as charge is consumed by stimulation current delivered by the DAC 60.
Further details concerning boost converter circuitry can be found in U.S. Pat. No. 7,872,884, which is incorporated herein by reference in its entirety. Moreover, one skilled in the art will realize that circuits other than a pulse width modulator 63 can be used in a boost converter. For example, a current- or voltage-controlled ring oscillator could also be used to toggle transistor 64.
While the boost converter 150 functions well to produce the desired compliance voltage V+, the inventors have noticed a shortcoming of such design. Specifically, the boost converter 150 has the potential to interfere with the telemetry circuitry operable in the IPG 100. FIG. 4 illustrates a typical bi-directional telemetry link operable between an IPG 100 and an external controller 12. As shown, the external controller 12 and the IPG 100 respectively contain transmitter/modulation and receiver/demodulation circuitry coupled to their coils 17 and 13 for communicating data between them. When data 170 is to be sent from the external controller 12 to the IPG 100, the data is modulated (e.g., encoded) and transmitted by circuitry 120 in the external controller. On the receiving side, this data 170 is received and demodulated (e.g., decoded) using circuitry 125 in the IPG 100. Similarly, when data 172 is to be sent from the IPG 100 to the external controller 12, the data is modulated and transmitted using circuitry 124 in the IPG. On the receiving side, this data 172 is received and demodulated using circuitry 121 in the external controller 12. As mentioned above, one modulation protocol operable in the respective modulation and demodulation circuit blocks 120, 121, 124, and 125 is FSK, which can represent logic ‘0’s and ‘1’s with an appropriate frequency. For example, logic ‘1’ can be modulated with a 129 kHz carrier, while logic ‘0 can be modulated with a 121 kHz carrier. The inductor-capacitor (LC) tank circuits associated with these links are accordingly tuned to resonate at these frequencies, as is well known.
Unfortunately, the boost converter 150, which also comprises an LC circuit, will also generate a magnetic field 173 when it is enabled, in particular because of the magnetic field generated by the inductor 66. This magnetic field 173 can interfere with the telemetry transmission and reception at coil 13 in the IPG 100. Even if coil 13 has a high quality factor, and good out-of-band noise rejection, the magnetic field 173 may still have frequency components that are within the band of coil 13 (e.g., from 100 kHz to 150 kHz). Moreover, the frequencies components present in magnetic field 173 can have a large bandwidth, and are difficult to control because they depend on the required compliance voltage V+ that must be produced at any given time. Because the IPG 100 usually allows a wide range of stimulation settings to be programmed, the possibility of telemetry interference arising from operation of the boost converter 150 becomes a real possibility. If the interference is severe, telemetry may not be possible during times when the IPG 100 is generating a compliance voltage, i.e., during times that the IPG 100 is operational and producing therapy to the patient, which is unpractical.
Accordingly, the implantable stimulator art would benefit from improved DC-to-DC converter circuitry for adjustably boosting the battery voltage to the compliance voltage needed to provide power to the stimulating electrode(s), while minimizing the effects of magnetic noise that interferes with telemetry operation of the implantable stimulator. Embodiments of such a solution are provided herein.